Dual-energy (DE) radiography is a well-known technique which allows the calculation of either bone-only or soft tissue-only images. The potential and clinical effectiveness of DE radiography has been well documented. DE radiography requires the acquisition of two images in which the effective energy of the detected x-rays for each image differs. The desired calculation is achieved by subtracting one image from the other. For a more detailed explanation of dual energy radiography, see W. R. Brody et al., A method for selective tissue and bone visualization using dual energy scanned projection radiography, Medical Physics, Vol. 8(3), May/June 1981, pps. 353-357. Presently there are two practical acquisition scenarios for dual energy imaging.
In one category of DE image acquisition, two images are acquired sequentially using a two-pulse method, wherein the kilovoltage and/or filtration is changed between the two pulses of x-rays, thereby changing the energy of the two images. Because of the problem of patient motion, the two images need to be acquired in rapid succession. This mandates a near real-time detector system such as an image intensifier (II) digital video system, as well as an x-ray generator capable of rapid kV switching. Van Lysel has recently reported on this type of dual energy acquisition for dual energy substraction in cardiac angiography, following earlier studies by others. While useful for many angiographic applications, II based imaging cannot meet either field of view or resolution requirements of general radiography.
Another category of dual energy acquisition utilizes a single x-ray pulse and stacked detectors. See D. L. Ergun et al., Single-Exposure Dual-Energy Computed Radiography: Improved Detection and Processing, Radiology, Medical Physics, Vol. 174, No. 1, 1990, pps. 243-249. The polyenergetic nature of the x-ray spectrum is exploited where a stack of detectors (with perhaps intervening filters) is used such that the first detector captures the lower energy image, and as the x-ray beam passes though the first detector and intervening filtration it is hardened before striking the second detector, which captures the higher energy image.
The single pulse used in this technique is attractive because the potential for motion artifacts is reduced. However, because the detectors are necessarily stacked they need to be very thin and thus only film/screen systems or stimulable phosphor systems can be used. Because stimulable phosphor systems to date are composed of only a single phosphor, typically BaFBr:Eu, differences in the so-called k-edge absorption properties cannot be exploited. In order to digitize the images from these stacked detector systems, they need to be separated and in the process the spatial alignment between the detectors is lost. This requires the re-registration of the images on a case by case basis.
While filters with different k-edges can be placed in a stacked detector between the fore and aft screens, the aft screen achieves energy separation (relative to the fore screen) only through the removal of x-ray photons from the x-ray beam, i.e, through filtration. Consequently, the high energy image derived from the aft screen will generally suffer from x-ray quantum noise effects, relative to the low energy image. Since the signal to noise ratio (SNR) in the subtracted image is maximized when the number of absorbed x-ray quanta are approximately equally distributed between the low energy and high energy images, true optimization is hard to achieve using the stacked stimulable phosphor approach.
As indicated previously, the stacked stimulable phosphor cassette must be dismantled in order to read the latent images on each of the screens using a scanning laser, requiring the images to be re-registered in the computer. Although correlation techniques have found to be a viable approach to the re-registration, these techniques are numerically intensive and could hinder practical clinical implementation of the technology. Furthermore, the scanning laser system employed to read out a stimulable phosphor screen is subject to time jitter noise similar to those in video systems. Since CCD cameras employ an array of discrete stationary detectors, time jitter noise does not exist for these devices. For more general background on the use of CCD cameras in radiography, see J.M. Heron et al., x-ray imaging with two-dimensional charge-coupled device (CCD) arrays, SPIE Vol. 486 Medical Imaging and Instrumentation '84, (1984), pps. 141-145.
Consequently, a need still exists for a dual energy radiography system which does not suffer from x-ray quantum noise effects, which does not require dismantling in order to analyze images and which is capable of real time imaging.